The present invention relates to cochlear prosthesis used to electrically stimulate the auditory nerve, and more particularly to a process for mapping a signal level into a stimulation current level.
Hearing loss, which may be due to many different causes, is generally of two types: conductive and sensorineural. Of these, conductive hearing loss occurs where the normal mechanical pathways for sound to reach the hair cells in the cochlea are impeded, for example, by damage to the ossicles. Conductive hearing loss may often be helped by use of conventional hearing aids, which amplify sound so that acoustic information does reach the cochlea and the hair cells. Some types of conductive hearing loss are also amenable to alleviation by surgical procedures.
In many people who are profoundly deaf, however, the reason for their deafness is sensorineural hearing loss. This type of hearing loss is due to the absence or the destruction of the hair cells in the cochlea which are needed to transduce acoustic signals into auditory nerve impulses. These people are unable to derive any benefit from conventional hearing aid systems, no matter how loud the acoustic stimulus is made, because their mechanisms for transducing sound energy into auditory nerve impulses have been damaged. Thus, in the absence of properly functioning hair cells, there is no way auditory nerve impulses can be generated directly from sounds.
To overcome sensorineural deafness, numerous implantable cochlear stimulation systems—or cochlear prosthesis—have been developed which seek to bypass the hair cells in the cochlear (the hair cells are located in the vicinity of the radially outer wall of the cochlea) by presenting electrical stimulation to the auditory nerve fibers directly, leading to the perception of sound in the brain and an at least partial restoration of hearing function. The common denominators in most of these cochlear prosthesis systems have been the implantation, into the cochlea, of electrodes, and a suitable external source of an electrical signal for the electrodes.
A cochlear prosthesis operates by direct electrical stimulation of the auditory nerve cells, bypassing the defective cochlear hair cells that normally transduce acoustic energy into electrical activity in such nerve cells. In order to effectively stimulate the nerve cells, the electronic circuitry and the electrode array of the cochlear prosthesis perform the function of separating the acoustic signal into a number of parallel channels of information, each representing the intensity of a narrow band of frequencies within the acoustic spectrum. Ideally, the electrode array would convey each channel of information selectively to the subset of auditory nerve cells that normally transmitted information about that frequency band to the brain. Those nerve cells are arranged in an orderly tonotopic sequence, from high frequencies at the basal end of the cochlear spiral to progressively lower frequencies towards the apex, and ideally the entire length of the cochlea would be stimulated to provide a full frequency range of hearing. In practice, this ideal is not achieved, because of the anatomy of the cochlea which decreases in diameter from the base to the apex, and exhibits variations between patients. Because of these difficulties, known electrodes can only be promoted to the second turn of the cochlea at best.
The signal provided to the electrode array is generated by a signal processing component of the Implantable Cochlear Stimulation (ICS) system. In known ICS systems, the acoustic signal is first processed by a family of parallel bandpass filters. Next the output of each bandpass filter is independently amplitude mapped into a simulation level using a mapping consistent with normal perception. In known systems, the mapping is a compressive mapping that is based on the log of the magnitude of each independent sample of the outputs of the band pass filters. The log is taken of the magnitude of each sample, then multiplied by a first scalar and added to a second scalar, and the sign of each sample is then applied to the compressed value. Disadvantageously, the log function can result in a DC component in the resulting signal, distorts sinusoidal inputs, and is computationally intensive.
The DC component arises from the asymmetry of the input waveform. The signal is processed before the amplitude mapping to remove DC bias, and as a result the total area under the waveform, at the output of the bandpass filters, sums to zero. But, the compressive nature of the log function reduces narrow high peaks much more than wide low peaks, and thereby creates a DC bias. A wideband speech signal is very asymmetric by nature, so the likelihood of generating such a DC bias is high. The presence of the DC bias poses a potential for tissue damage after long term use, and may cause the charge in a capacitor typically, used for energy storage in the implantable stimulation circuit, to grow large resulting in undesirable nonlinear behavior.
The shape of a waveform processed by the amplitude mapping may be distorted by the compression. For example, samples from the peak of a sinusoidal waveform are compressed more than samples between the peaks, and as a result the sinusoid becomes more like a square wave with rounded corners than like a sinusoid. When patients are tested for psychophysical thresholds, sine waves are used as the stimulating signals for each electrode. The frequency of each sine wave is selected as the center frequency of the band pass filter that processes the signal for the corresponding electrode in normal system operation. When the threshold levels determined during psychophysical testing are later applied to a compressed sinusoid, which compressed sinusoid has the same peak stimulating current as the original sinusoid that the thresholds are based on, the perceived loudness may not be the same as with the original sinusoid. Although the peak stimulation currents of the original sinusoid and the compressed sinusoid are the same, the amplitude mapping brings up the “shoulders” on the compressed sinusoid, making it more like a square wave with rounded corners. As a result of “raising the shoulders” of the sinusoid, charge per phase raises, which results in the perceived loudness increasing. This increase in perceived loudness may be significant for patients with a narrow dynamic hearing range.
The processing required to compute the log of each sample, in each frequency band, at a high data rate, is a computationally demanding process that expends significant power in the signal processor. The development of Behind-The-Ear (BTE) speech processor, and fully implantable cochlear stimulators, requires that power consumption be reduced to a minimum. A BTE ICS system is described in U.S. Pat. No. 5,824,022 issued Oct. 20, 1998 for ‘Cochlear stimulation system employing behind-the-ear speech processor with remote control.’ Behind-the-ear speech processors offer several advantages, but their small size limits the size of the battery they may carry (which in turn limits the capacity of the battery.) The small battery size results in a requirement for very low power consumption. Processing, such as that required by known amplitude mapping methods, work against the need to reduce power dissipation. The '022 patent is herein incorporated by reference.
An improvement to the current compressive processing is needed to both improve performance, and to reduce the power consumption required for signal processing.